REVIEW PAPER

Recent progress on magnetic nanoparticles for magnetichyperthermia

Lina Kafrouni1,2 • Oumarou Savadogo1,2

Received: 23 February 2016 / Accepted: 19 August 2016 / Published online: 6 September 2016

� The Author(s) 2016. This article is published with open access at Springerlink.com

Abstract Recent advances in nanomaterials science con-

tributed to develop new micro- and nano-devices as

potential diagnostic and therapeutic tools in the field of

oncology. The synthesis of superparamagnetic nanoparti-

cles (SPMNPs) has been intensively studied, and the use of

these particles in magnetic hyperthermia therapy has

demonstrated successes in treatment of cancer. However,

some physical limitations have been found to impact the

heating efficiency required to kill cancer cells. Moreover,

the bio-safety of NPs remains largely unexplored. The

primary goals of this review are to summarize the recent

progress in the development of magnetic nanoparticles

(MNPs) for hyperthermia, and discuss the limitations and

advances in the synthesis of these particles. Based on this

knowledge, new perspectives on development of new

biocompatible and biofunctional nanomaterials for mag-

netic hyperthermia are discussed.

Keywords Magnetic nanoparticles � Synthesis � Magnetic

hyperthermia � Cancer

Introduction

According to the National Cancer Institute, cancer is cur-

rently the second leading cause of death in the United

States, exceeded only by heart disease as the number one

killer. A total of 1,620 Americans are expected to die of

cancer per day in 2015.

Significant progress has been made so far in nanotech-

nology for the diagnosis and treatment of cancer. A variety

of magnetic nanomaterials has been developed to achieve

improved efficacy in cancer therapy as well as reduced side

effects compared to conventional therapies. The interest in

MNPs is due to their unique magnetic properties; they

exhibit diagnostic tool, drug carrier and heat generator for

therapy in magnetic resonance imaging (MRI), so-called

‘theranostic’ and their small sizes, which allow the parti-

cles to reach most biological tissues. Currently, iron oxide

nanoparticles (IONPs) are the most explored MNPs for

magnetic hyperthermia, because of their lack of toxicity

and their known pathways of metabolism (Tran et al.

2012a, b).

The generation of heat by the exposition of MNPs to a

non-invasive alternating magnetic field (AMF) can be used

to destroy tumor tissue, given that heat promotes cell

apoptosis through irreversible physiological changes (Pra-

sad et al. 2007). This approach is known as magnetic

hyperthermia. The basics of the magnetic properties

required in MNPs for magnetic hyperthermia applications

will be discussed later in detail.

The synthesis methods of MNPs have an impressive

impact on the magnetic and morphological properties of

the final product (Castellanos-Rubio et al. 2015). There-

fore, a synthesis method with the ability to rigorously

control the composition, size and shape is needed. This

paper presents a short review on the current methods for

& Oumarou Savadogo

osavadogo@polymtl.ca

1 Department of Chemical Engineering, Polytechnique

Montreal, C.P. 6079, Succursale Centre-ville, Montreal,

QC H3C 3A7, Canada

2 Laboratory of New Materials for Energy and

Electrochemistry Systems (LaNoMat), Montreal, Canada

123

Prog Biomater (2016) 5:147–160

DOI 10.1007/s40204-016-0054-6

synthesis of MNPs for nanomedicine, and discusses

important findings reported earlier.

Basics of magnetism in magnetic hyperthermia

An understanding of the relationship between physico-

chemical properties (for example: structure, particle size)

and magnetic properties is essential to design new mag-

netic materials for magnetic hyperthermia applications.

Therefore, a review on the basic concepts in nano-mag-

netism will be discussed shortly.

Soft and hard magnets

When a ferromagnetic material, such as Iron, nickel and

cobalt, is placed in a magnetic field of strength ‘H’, the

atoms acquire an induced magnetic moment ‘m’ randomly

oriented. The magnetic moments pointed in the same

direction per volume of atoms are called magnetization

‘M’. The magnetic induction ‘B’ is given by Maxwell’s

equation (Eq. 1) (Laurent et al. 2011).

B ¼ l0 H þMð Þ ð1Þ

where l0 is the permeability of the free space which equals

to 4p 10�7 V.s/A.m.

The small regions of magnetization are called magnetic

domains, and the boundaries between domains are called

domain walls. In the absence of an external magnetic field,

ferromagnetic material does not show any magnetization

due to the random orientation of the magnetizations in

magnetic domains (Point a, Fig. 1). However, when an

external magnetic field is applied, magnetic moments

become aligned to the direction of the magnetic field, so

the domain walls disappear and the magnetization becomes

saturated, the so-called saturation magnetization (Ms)

(Fig. 1).

Once the applied magnetic field is removed, ferromag-

netic materials keep some memory of the applied field

(Point b, Fig. 1), called remanence (Mr). A coercive force

must be applied to reduce the remanent magnetization to

zero and close the loop.

Ferromagnetic materials can be categorized into soft and

hard magnets (Mody et al. 2013). Soft magnets have a low

coercivity (Hc), so they can be demagnetized at low

magnetic field. However, hard magnets exhibit a high Hc

and thus they are difficult to demagnetize.

Multi-domain to single domain

The magnetostatic (dipole–dipole) energy is inversely

proportional to the volume of the particle (r3), and the

domain-wall energy is proportional to the area of the wall

(r2) (Fig. 2) (Spaldin 2011).

By looking at the balance between the magnetostatic

energy and the domain wall energy, it is energetically

unfavorable to form domain walls below a critical radius,

because the domain-wall energy is very low, and a single

domain is formed as a result of high magnetostatic energy.

For a sphere containing two semi-sphere domains of

opposite magnetization with axial magnetic anisotropy, the

critical single-domain radius is given by Eq. (2) (Skomski

2003).

rcritical ¼36

ffiffiffiffiffiffiffiffiffi

AK1

p

l0M2s

ð2Þ

Fig. 1 Typical hysteresis loop of ferromagnetic materials (adapted

from Mody et al. 2013)

Fig. 2 Relative stability of multi-domain and single domain (adapted

from Spaldin 2011)

148 Prog Biomater (2016) 5:147–160

123

where A is the exchange stiffness and K1 is the first uniaxial

anisotropy constant.

The critical radius values corresponding to ferromag-

netic elements Fe, Co and Ni are calculated according to

Eq. (2) and are presented in Table 1.

Superparamagnetism

It has been found that with a further decrease in particle

size below the critical radius, the coercivity Hc decreases

significantly to reach zero. When the coercivity becomes

zero, the particles magnetize in the presence of an external

magnetic field and revert to a non-magnetic state when the

external magnetic field is removed (Fig. 3) (Mody et al.

2013).

This behavior can be explained by the fact that a small

magnetic particle less than critical size prefers to be uni-

formly magnetized along one of its easy axes (h = 0,

h = p), and the energy required to rotate the magnetization

away from the easy direction is called magnetic anisotropy

energy. In a simple model for a non-interacting single-

domain spherical particle with uniaxial anisotropy in zero

magnetic field, the magnetic anisotropy energy ‘EA’ is

given by an expression of Eq. (3) (Stoner and Wohlfarth

1948).

EA ¼ K � V � sin2 h ð3Þ

where K is the anisotropy constant, V is the volume of the

particle and h is the angle between the particle magneti-

zation and the easy magnetization axis of the particle.

According to Eq. (3), the magnetic anisotropy energy

decreases when the volume of the particle becomes smal-

ler. Furthermore, the anisotropy energy becomes compa-

rable to or even lower than the thermal energy

(Ethermal = kB�T, where kB is Boltzmann constant) (Krish-

nan 2010). As a result, the energy barrier for magnetization

reversal can be overcome thermally (Fig. 4). This phe-

nomenon is called ‘superparamagnetism’.

Due to the fact that these particles are magnetically

controlled by an external magnetic field and maintain a

colloidal stability upon removal of the external magnetic

field, superparamagnetic particles have a unique advantage

for biomedical applications.

For spherical magnetic particles, the transition from

single domain to superparamagnetic ‘r0’ depends on the

size and/or geometry of the particles and can be determined

by the following Eq. (4) (Martel 2015):

r0 ¼6 � kB � TB

K

� �1=3

ð4Þ

where TB is the blocking temperature.

Table 2 provides calculated values of the transition

radius ‘r0’, according to Eq. (4), for the main magnetic

nanomaterials (Martel 2015; Kolhatkar et al. 2013).

Although particle moves toward superparamagnetism

when the size of the particle decreases below the transition

point and becomes suitable for biomedical application, the

saturation magnetization ‘Ms’ reduces. The magnitude of

Table 1 Magnetic parameters at room temperature (Skomski 2003)

Ferromagnetic particles Fe Co Ni

A (pJ/m) 8.3 10.3 3.4

K1 (MJ/m3) 0.05 0.53 -0.005

l0�Ms (T) 2.15 1.76 0.61

rC (nm) 6 34 16

Fig. 3 The magnetic response characteristic of a superparamagnetic

material (adapted from Mody et al. 2013)

Fig. 4 Schematic of anisotropy energy barrier for magnetization

reversal (adapted from Stoner and Wohlfarth 1948)

Prog Biomater (2016) 5:147–160 149

123

Ms is inversely proportional to the ratio of disordered spin

layer at the surface to the radius of the particle, which

significantly increases when the size of the nanoparticle

becomes too small. The relationship between Ms, the size

and the disordered spin layer is described by Eq. (5) (Jun

et al. 2008):

Ms ¼ Msbðr � dÞ

r

� �3

ð5Þ

where d is the thickness of the particle’s surface exhibiting

disordered spins, and Msb is the bulk Ms. Recent studies on

the effect of the size of MNPs upon its saturation magne-

tization are summarized in Table 3. According to the

studies listed in Table 3, the Ms increases with the size of

the MNPs due to the reduction of the spin disorder effect.

Recent study done by Guardia et al. have demonstrated

that the surface coating of iron oxide (Fe3O4) NPs with

oleic acid increases their measured Ms to reach the bulk

value, by reducing the level of surface spin disorder

(Guardia et al. 2007).

Heat generation

Heating tumor cells with SPMNPs by magnetic hyper-

thermia is based on Neel and Brownian relaxations. In the

presence of an external alternating magnetic field, the

magnetic moment rotates and the nanoparticle itself

rotates, then relaxes back to their original magnetic field

orientation. The rotation of the magnetic moment (Neel

mode) and the friction arising from particle oscillations

(Brownian mode) leads to a phase lag between applied

magnetic field and the direction of the magnetic moments.

As a result, the heat is released.

The efficiency of heating is measured in terms of the

specific absorption rate (SAR), or specific loss of power

(SLP), which is defined in Eq. (6). For biomedical appli-

cations, the value of SAR is crucial, because the higher the

specific absorption rate, the lower the injected dose to the

patient.

SAR or SLP W=gð Þ ¼ CDTDt

ð6Þ

where C is the specific heat capacity of water, and DT/Dt isthe rate of change of temperature versus time.

According to Rosensweig (2002), there is a strong

relationship between the SAR of SPMNPs and its magnetic

relaxation ‘s’ (Eq. 7).

SAR ¼ 4:1868pl20uM2

s V

1000kT� H2

0v2pms

1þ ð2pmsÞ2ð7Þ

where u is the volume fraction of the SPMNPs, V = 4pr33

is

the magnetic volume for a particle of radius r, H0 is the

magnetic field intensity, m is the frequency of the oscillat-

ing magnetic field and s is the relaxation time. The other

parameters (lo is the permeability of the free space), p,k (Boltzmann constant) and T (temperature of the sample)

have their classical meanings.

Also, Eq. (7) shows that the SAR strongly depends on

the Ms and the volume fraction of the SPMNPs. Not only

high Ms values are required for thermal energy dissipation

in the tumor cells, but also to give more control on the

magnetophoretic velocity of the MNPs ‘Vmag’ in the blood

using external magnetic field (Grief and Richardson 2005)

(Eq. 8).

Vmag ¼Ms � Vmicrodevice � rB

6 � p � Rmicrodevice � lð8Þ

where Vmicrodevice is the volume of microdevice (m3), rB is

the magnetic gradient applied (T/m), Rmicrodevice is the

microdevice radius (m) and l is the blood viscosity (Pa.s).

Theoretically, a critical diameter dc is defined as the

diameter for which the Neel relaxation time ‘sN’ (Eq. 9) isequal to the Brownian relaxation time ‘sB’ (Shliomis and

Stepanov 1990) (Eq. 10). For small particles with a

diameter\dc, Neel relaxation is predominant. However,

the heating is primarily due to Brownian rotation in larger

particles with a diameter[dc. The dominating contribution

will be by the faster relaxation time.

sN ¼ s0eK�VkB �T ð9Þ

Table 2 Maximum radius for superparamagnetic NPs of different compositions (Martel 2015; Kolhatkar et al. 2013)

Superparamagnetic NPs Co CoPt Co Fe2O4 FeCo Fe3O4 Fe2O3 FePt Ni

r0 (nm) 5 1 5 10 12.5 15 1.5 15

Table 3 Magnetizations of a variety of types of MNPs of varying

sizes

MNPs Size (nm) Ms (emu/g) References

Co Fe2O4 4.2 30.6 Pereira et al. (2012)

4.8 46.0

18.6 48.8

Fe3O4 4.9 60.4

6.3 64.8

Ni 24 25.3 He and Shi (2012)

50 32.3

150 Prog Biomater (2016) 5:147–160

123

sB ¼ 3gVB

kB � Tð10Þ

where K is the anisotropy constant of magnetite which is

over the range of 23,000–100,000 J m-3, while

s0 & 10�9–10�12 s is the relaxation time of non-interact-

ing MNPs, g is the viscosity of the surrounding liquid and

VB is the hydrodynamic volume of the particle; kB is the

Boltzmann constant and T is the temperature of the sample.

The frequency mN for maximal heating via Neel relax-

ation is given by Eq. (11), and the frequency mB for max-

imal heating via Brown rotation is given by Eq. (12)

(Fannin and Charles 1991).

2pmNsN ¼ 1 ð11Þ2pmBsB ¼ 1 ð12Þ

When the diameter of the particle is close to dc,

sN & sB and an effective relaxation time ‘seff’ is defined inEq. (13). The frequency for maximal heating ‘meff’ is thengiven by Eq. (14) (Fannin et al. 1993).

seff ¼sNsB

ðsN þ sBÞð13Þ

2pmeffseff ¼ 1 ð14Þ

Recent research optimized the heating efficiency by

tuning the MNPs size to match the total relaxation time

(stotal = sN ? sB) to the applied frequency (m = 12p�stotal)

(Khandhar et al. 2011).

The strong dependence of the SAR on multiple magnetic

properties such as saturation magnetization and relaxation

time, physical parameters like size, shape and composition

can be tailored to enhance the heat dissipation and thus

lower the injected dose of SPMNPs in the tumor site.

Biomaterials for magnetic hyperthermia

To develop excellent candidates for magnetic hyperther-

mia, it is very important to review the recent advances and

limitations in the development of MNPs for magnetic

hyperthermia applications.

Superparamagnetic iron oxide nanoparticles (SPIONs)

are the most used MNPs for biomedical applications,

especially magnetic hyperthermia. They received consid-

erable attention due to their biocompatibility compared to

other magnetic materials such as cobalt and nickel (Tran

et al. 2012a, b). The high biocompatibility of IONPs is due

to well-controlled cell homeostasis by uptake, excretion

and storage (Chenga et al. 2005).

However, nickel and cobalt are susceptible to oxidation

and toxic, even though they exhibit a high magnetic

moment, because they are not essential elements to the

body like iron and thus accumulate in the body and cause

illness. It is worth noting that SPIONs may induce dermal

toxicity via their ability to be internalized and thereby

initiate oxidative stress leading to inflammation (Murray

et al. 2013).

IONPs become superparamagnetic at room temperature

when their radius is below about 15 nm (Kolhatkar et al.

2013), and aggregation is a common phenomenon among

SPIONs (Wu et al. 2015). Therefore, bare SPIONs are

coated against aggregation by either non-magnetic or

magnetic shell (Zenga et al. 2004). Usually, the type of

coatings has an impact on the heating efficiency of the core

through modifying the surface properties. Details on the

types of shells used to protect IONPs and their effect over

magnetic properties will be discussed.

Among iron oxides, magnetite (Fe3O4) and maghemite

(c-Fe2O3) are very popular candidates and have unique

magnetic properties suitable for biomedical applications.

Iron metal (Fe) has a higher magnetization than mag-

netite and maghemite. However, Fe is highly susceptible to

oxidation, which limits its use for biomedical applications.

Qiang et al. synthesize oxidative stable Fe-core MNPs

coated with iron oxide and having an increasing Ms from

about 80 emu/g (at the cluster size of 3 nm) to 200 emu/g

(at the size of 100 nm) (Qiang et al. 2006).

In general, MNPs are coated with a selected material to

enhance their colloidal stability and biocompatibility or to

offer them the capacity to functionalize the surface, like in

the case of a coating of silica (SiO2) (Rittikulsittichai et al.

2013). Furthermore, coating can be used to modify MNPs

surface to increase their Ms and consequently increase the

SAR.

Studies show that coating MNPs with non-magnetic

material, for example Fe3O4 coated with SiO2 (Larumbe

et al. 2012a, b), will reduce Ms (from 72 emu/g to 37 emu/

g) and hence SAR (from 1.5 ± 0.1 to 1.08 ± 0.04 W/g) as

compared to uncoated MNPs. The decrease in Ms was

attributed to the enhanced surface spin effects, and thus not

all the IONPs mass contribute to Ms. Furthermore, the

effective anisotropy constant ‘Keff’ increases due to the

strain and surface spin disorders created by SiO2 coating,

and the blocking temperature TB experiences similar vari-

ations since TB is defined as the product of the Keff and the

volume of the nanoparticles ‘V’ (Eq. 15) (Coskun et al.

2010).

TB ¼ Keff � V25 � kB

ð15Þ

Surface spin effect (or surface spin disorder) is the result

of the surface electrons engagement in the bond with the

coating material, which no longer participate in the mag-

netic super-exchange bonds between metal cations (ex-

ample: Fe–O–Fe), and thus reduce the coordination

between surface spins (Kodama et al. 1996).

Prog Biomater (2016) 5:147–160 151

123

Fe3O4 NPs coated with SiO2 and functionalized with

propylamine groups showed higher magnetization satura-

tion (Ms & 42 emu/g) than uncoated Fe3O4

(Ms & 27 emu/g), where both were synthesized by ther-

mal decomposition in oleic acid (Woo et al. 2005). It seems

that the surface of Fe3O4 is magnetically more active in

Fe3O4 NPs coated with silica-propylamine than that of

uncoated Fe3O4 covered with oleic acid.

On the contrary, Fe3O4 NPs coated with silica-propy-

lamine showed slightly lower magnetization saturation

(Ms & 58 emu/g) than uncoated Fe3O4 (Ms & 60 emu/g)

(Yamaura et al. 2004), where Fe3O4 NPs were obtained by

co-precipitation in aqueous medium. The contradictory

results of these two studies suggest that the synthesis and

coating methods can be tailored to enhance the magnetic

properties of the MNPs.

Capping Co-MNPs with metallic shell (such as Cu or

Au) provides us a high tuning opportunity over the mag-

netic properties (for example, enhanced surface anisotropy

and higher blocking temperature), due to the bonding of the

d-orbital electrons of the core to the conduction band

orbitals of the capping layer (Luis et al. 2006). This sug-

gests that the surface anisotropy is mainly determined by

the electronic states of the core–shell metals and, therefore,

it could be tuned by choosing materials with appropriate

electronic band structures.

For hyperthermia applications, an SLP of 1000 W/g is

necessary at 100 kHz and 20 mT (human-compatible

conditions). By taking advantage of the exchange coupling

between a magnetically hard core (CoFe2O4) and soft shell

(Mn Fe2O4), MNPs exhibiting a significant enhancement in

SLP have been developed (Lee et al. 2011). Various

combinations of core–shell nanoparticles tuned Ms of the

single-component MNPs to achieve high SLP while

maintaining the superparamagnetism. For example

Zn0:4Co0:6Fe2O4 core and Zn0:4Mn0:6Fe2O4 shell MNPs

have an SLP of 3866 W/g and thus exhibit 1.7 times higher

SLP than that for CoFe2O4(core) MnFe2O4(shell) MNPs

(2274.12 W/g) and 34 times larger than that for commer-

cial Feridex Fe3O4 NPs (114 W/g).

Spherical Mn Fe2O4 SPMNPs show lower SLP of

411 W/g (r = 15 nm) when compared to that of Mn

Fe2O4(core) Co Fe2O4(shell) (r = 15 nm) where SLP is

about 3034 W/g (Noh et al. 2012). Clearly, core–shell

design has the advantage in achieving large SLP while

keeping the superparamagnetism of the nanoparticle. In the

same work, cubes of Co Fe2O4 coated with Zn0:4Fe2:6O4

showed a 4-fold increase in coercivity as compared to the

core alone. This increase is consequently followed by a

dramatically higher SAR for the shell-core MNPs

(10,600 W/g) when compared to that of MNPs composed

of just the core (4060 W/g).

Many efforts have been dedicated toward understanding

the relationship between the shape of MNPs and their

magnetic properties. Several studies showed that the Ms is

proportional to the volume of the particle (V) with the same

crystalline composition but different shape (Chou et al.

2009; Shevchenko et al. 2003), due to the decrease of the

surface-to-volume ratio and consequently surface spin

disorder. For example, considering MNPs having the same

unit size (d) (where ‘d’ corresponds to the side length for

nanocubes, the width for nanorods and the diameter for

nanospheres), the V of nanocube is higher than the V of

nanorod, and the V of nanosphere is lower than the V of

nanorod. Therefore, the same order of Ms is expected (Ms

of nanocube[Ms of nanorod[Ms of nanosphere).

A study on the effect of the shape of Fe3O4 NPs over its

saturation magnetization is done by Zhen et al. (Zhen et al.

2011). The authors observed a higher Ms for the cubic

shape (Ms = 40 emu/g) compared to the spherical shape

(Ms = 31 emu/g), where the volume of the cube is slightly

higher than that of the sphere (Vcube [Vsphere). They

attributed the lower magnetization of spherical Fe3O4 NPs

to their crystalline defect structure or greater degree of

oxidation and non-magnetic iron oxide (Fe2O3) content.

According to Noh et al. (2012), the cubic shape of

Zn0:4Fe2:6O4 has a higher Ms (165 emu/g) value than the

spherical shape (145 emu/g) with the same volume. In fact,

the surface of the cube shape has a smaller surface aniso-

tropy since its topology comprises low energy facets. As a

result, disordered magnetic spins in cubic NPs (4 %) are

lower than in spherical NPs (8 %).

However, in a study done by Montferrand et al. on

Fe3O4 NPs (Montferrand et al. 2013) Ms for the cubic

shape (40 emu/g) is lower than the spherical shape

(80 emu/g) of the same size. Unexpected Ms could be

related to size polydispersity and polymorphism detected in

TEM images.

Magnetic properties are also defined by the atomic state

of the elements, especially the number of unpaired valence

electrons. For example, Fe(III) have five unpaired electrons

and thus a moment of 5 9 1.73 = 8.65 Bohr magnetons.

Moreover, the distribution of ions in the structure is another

parameter responsible for the determination of the moment.

For example, in an inverse spinel structure of ferrites, the

magnetic moments of the cations in the octahedral sites are

aligned parallel to the magnetic field, and the ones in the

tetrahedral sites are antiparallel, leading to a decrease in the

net moment (Lee et al. 2006).

Hence, doping MNPs with cations is of great interest in

nanomedicine because it tailors the physical and magnetic

properties, without affecting its crystal structure, due to the

nature of the cation and its relative distribution in the

tetrahedral and octahedral sites (Fantechi et al. 2012).

152 Prog Biomater (2016) 5:147–160

123

Lee et al. (2006) compared the crystal structure of four

spinel ferrites (M Fe2O4): MnFe2O4 (110 emu/g), Fe

Fe2O4 (101 emu/g), CoFe2O4 (99 emu/g), and NiFe2O4

MNPs (85 emu/g). MnFe2O4 had a mixed spinel structure,

where Mn2þ and Fe3þ occupied both octahedral and

tetrahedral sites, and an inverse spinel structure where

Mn2þ and Fe3þ occupied octahedral sites and only Fe3þ

occupied the tetrahedral sites.

The inclusion of Ni2þ in the ferrite spinel structure

(NixFe3�xO4 with x = 0, 0.04, 0.06 and 0.11) has no

substantial change in the value of Ms, where Ni2þ occupy

Fe2þ octahedral sites (Larumbe et al. 2012a, b). Gabal et al.

examined the Zn2þ doped nickel ferrite (Ni1�xZnxFe2O4;

0\ x\ 1) and noticed that the Ms increases by increasing

Zn doping levels up to 0.5 (Jalalya et al. 2010). This

behavior can be explained by the fact that magnetite

(Fe3O4), with a spinel structure, has Fe3þ ions occupying

tetrahedral (inverse) sites and Fe2þ with Fe3þ ions residing

in the octahedral sites. During cation exchange Fe2þ in

octahedral site is replaced by Ni2þ and NiFe2O4 is formed.

Since the tetrahedral and octahedral sites are antiferro-

magnetically coupled, the net moment of Ni ferrite equals

the moment of octahedral site (Ni2þ, Fe3þ) minus the

moment of tetrahedral (Fe3þ). The inclusion of non-mag-

netic Zn2þ in NiFe2O4 substitutes Ni2þ then occupies a

tetrahedral site and force magnetic Fe3þ to migrate to

octahedral site and, as x increases. As a result, the net

moment increases due to the decrease in fraction of

moment of tetrahedral site and an increase in the moment

of octahedral sites (Jalalya et al. 2010).

FeCo MNPs usually exhibit high Ms values (122–

230 emu/g) compared with CoFe2O4 MNPs (Chaubey

et al. 2007), due to the absence of the non-magnetic oxygen

component (Zhang et al. 2012). However, the ease of

oxidation in the presence of air is the key issue for these

alloys (Zhang et al. 2012).

Palladium metal is a non-magnetic element, but tends to

order ferromagnetically when alloyed with a small amount

of magnetic transition metal impurities (such as Fe, Co and

Ni 3d metals) (Crangle and Scott 1965). A polarization of

Pd atom by a magnetic impurity is due to the hybridization

and exchange between 4d and 3d orbitals (Fig. 5) (Van

Acker et al. 1991).

The appearance of ferromagnetism can be explained by

the large density of states at the Fermi level (EF).

Paulus and Tucker (1995) proposed for the first time

PdCo seeds for thermal treatment of tumors. PdCo ther-

moseeds (typically rod shape where d = 1 mm and

L = 1–2 cm) are permanently implanted into the cancerous

tissue, and thus the patient can be scheduled for activation

of the thermoseeds at intervals of minimal thermotolerance

(Paulus and Tucker 1995). The authors developed a new

approach to treat prostate cancer, post-radiotherapy, using

these thermoseeds. During thermotherapy, PdCo rods heat

up when exposed to an alternative magnetic field (due to

eddy current) to a specific temperature (Curie temperature),

at which the alloy goes from being magnetic to non-mag-

netic, and ceases to heat up and it simply maintains the

Curie temperature as long as it remains in the magnetic

field (Paulus and Tucker 1995).

Deger et al. (2002) conducted a clinical study on the

treatment of patients with localized prostate cancer with a

magnetic hyperthermia, using self-regulating PdCo ther-

moseeds, after radiotherapy. During hyperthermia, PdCo

thermoseeds heating temperatures were between 42 and

46 �C with a Curie temperature of 55 �C. The initial

median prostate-specific antigen (PSA) value was 11.6 ng/

ml, and then decreased to 1.3 and 0.55 ng/ml after 12 and

24 months, respectively, after the therapy. Moreover, PdCo

seeds proved to be biocompatible and do not show major

complication during the treatment, and remain in the

prostate during follow up (Deger et al. 2002).

According to Brezovich and Meredith (1989), a heat

production rate of 200 mW/cm is adequate for most clinical

application. El-Sayed et al. calculate the power dissipated

from Pd89:2Co10:8,Pd73Ni27 and Cu29:6Ni70:4 ferromagnetic

seeds, having a rod shape with a diameter of 0.9 mm

diameter and a 5.5 cm length as function of temperature

(El-Sayed et al. 2007). At 20 �C, the heating power of

Pd89:2Co10:8 was about 171 mW/g, and 150 mW/g for

Pd73Ni27. The Cu29:6Ni70:4 seed showed a much smaller

heating power of 80 mW/g. Therefore, Pd89:2Co10:8 seed

exhibited the highest heating power to treat localized

tumors compared with the other two alloys.

Iron based-MNPs have been widely studied for nano-

medicine (especially for cancer treatment) and palladium-

cobalt alloys have not received significant attention.

Although Pd and Co are toxic elements, PdCo alloy has a

higher stability and resistance to corrosion (Wataha et al.

1991) compared to Fe-based alloy (Arbab et al. 2005).

Moreover, the researches done over PdCo thermoseeds are

very promising and encouraging to develop new MNPs

candidates for thermotherapy made of PdCo alloys.

Fig. 5 Illustration of the covalent interaction between Fe 3d and Pd

4d orbitals (reproduced from Van Acker et al. 1991)

Prog Biomater (2016) 5:147–160 153

123

Nanotoxicity of biomaterials

Considering the wide preclinical and clinical applications

of magnetic iron oxide NPs in nanomedicine, it is crucial to

understand the potential nanotoxicity associated with

exposure to these NPs and especially the physiological

effects produced by the surface coatings used for

functionality.

The work of Pisanic et al. (2007) showed that the

intracellular delivery of 0.15–15 mm of iron oxide (Fe2O3)

NPs may adversely affect cell function and results in a

dose-dependent diminishing viability and capacity of PC12

cells (rat pheochromocytoma cell line) to differentiate, in

response to nerve growth factor.

In fact, uncoated iron oxide NPs have a low solubility

that can lead to their precipitation and a high rate of

agglomeration under physiological conditions (Lei et al.

2013). Coating these NPs aims to stabilize their surfaces

against agglomeration and dissolution, and allows the

grafting of biomolecules (such as antibodies and drugs)

(Sadeghiani et al. 2005). However, the type of surface-

coating materials is important to determine the toxicity of

coated NPs.

The cytotoxic potential of iron oxide NPs with a range

of surface coatings has been extensively investigated.

Hilger et al. (2003) estimated the cytotoxic potential of

cationic/anionic coated magnetite (Fe3O4) nanoparticles by

measuring the succinate dehydrogenase activity in human

adenocarcinoma cells (BT-20). Cationic particles showed

to induce the strongest decrease in cell survival rates of

BT-20 cells (0 ± 0 after incubation for 72 h) for a con-

centration of 20 ng/cell. This is due to some strong elec-

trostatic bindings to cellular membranes. On the other

hand, Berry et al. (2003) found that dextran-coated iron

oxide NPs could induce cell death and reduce proliferation

of human fibroblasts during internalization. Significant

membrane disruptions were observed in fibroblasts cells,

including possible apoptosis and aberrations in cell mor-

phology, causing decreases in cells motility (Berry et al.

2004).

Recent studies show that Fe3O4 NPs can affect the

cellular functionality by altering the level of transferrin

receptor expression and can change the cellular prolifera-

tion capacity by altering the expression of cyclins and

cyclin-dependent kinases in cell cycle (Schafer et al. 2007;

Huang et al. 2009). Moreover, researchers are finding

evidence that Fe3O4 NPs exposure can produce mutagenic

effects including: chromosomal aberrations, DNA strand

breakage, oxidative DNA damage and mutations (Koedrith

et al. 2014). Other research has reported that the excess of

iron exposure has been found to cause elevated ROS

generation through the Fenton reaction, resulting in

oxidative stress that damages DNA, lipids and proteins,

consequently resulting in carcinogenesis (Toyokuni 1996).

These findings confirm previous reports that the presence

of intracellular Fe3O4 nanoparticle constructs can result in

significant changes in cell behavior and viability

(Buyukhatipoglu and Clyne 2011).

Upon administration into tumor tissue, MNPs interact

with blood components, where thousands of biomolecules

compete for limited space on an NP surface (Cedervall

et al. 2007), due to van der Waal’s interactions, electro-

static interactions, hydrogen bonding and/or hydrophobic

interactions (Hlady and Buijs 1996). As a result, MNPs

acquire a dynamic exchange plasma proteins layer, so-

called ‘corona’ (Cedervall et al. 2007), in which competi-

tive displacement of earlier adsorbed proteins by other

proteins with stronger binding affinities takes place and is

referred to as ‘Vroman Effect’ (Hirsh et al. 2013). Thus, the

identity, organization and residence time of these proteins

determine the way cells interact with NPs (Cedervall et al.

2007). Moreover, the adsorbed proteins identity and their

total amount showed to be strongly dependent on the par-

ticle surface chemistry (like surface composition, charge,

topography and area) (Hlady and Buijs 1996).

Studies show that plasma proteins, including

immunoglobulins and complement proteins, once adsorbed

to NPs surfaces it target the particles as pathogens for

clearance (called ‘opsonization’) by the reticulo-endothe-

lial system and mononuclear phagocytic system (Ehren-

berg et al. 2009). In fact, the immune system may

recognize the proteins as native or as foreign pathogen

depending on whether the proteins bind or not to immune

cells receptors. Following proteins adsorption, platelets

cells adhesion and activation on NPs may occur via inter-

action of adhesion receptors with the adsorbed blood pro-

teins such as fibrinogen, fibronectin, vitronectin, and the

von Willebrand factor (Nygren et al. 1995; Elam and

Nygren 1992). As a result, inflammatory cells (primary

polymorphonuclear leukocytes) migrate from the blood

toward the NPs, triggered by chemoattractants released

from activated cells (Franz et al. 2011). Inflammatory cells’

adsorption over the protein-coated NPs surface, due to

protein ligands of integrins, leads to an acute or chronic

inflammation (Nimeri et al. 2002).

The concept of inert biomaterials points out the need of

strategies for improving implant integration, to avoid for-

eign body reactions. It was shown that when macrophages

are cultured on surface-modified polymers displaying

hydrophobic, hydrophilic and/or ionic chemistries, they

change their protein expression profiles and cytokine/che-

mokine responses (Dinnes et al. 2007). Consequently,

current studies in the design of such biomaterials include

passive modulation of the surface chemistry, to limit

154 Prog Biomater (2016) 5:147–160

123

immune responses. For example, polyethylene glycol

(PEG)-modified surface reduces protein adsorption due to

its sterically hindered and hydrophilic coating (Torchilin

and Papisov 1994), and this leads to more blood circulation

of PEG-coated NPs. On the other side, functionalization of

the surface with bioactive molecule such as adhesion sites

(Kao and Lee 2001), anti-inflammatory drugs (Franchi-

mont et al. 2002) and growth factors (Barrientos et al.

2008) is also a very interesting strategy for modulating or

suppressing inflammatory responses.

MNPs can induce toxicity, not only by activating cells in

a direct way as discussed above, but also indirectly by

excessive tissue accumulation of free metal ions (Weir

et al. 1984). It was shown that reactive oxygen species

(ROS) are generated by the cells as a result of leached ions

after exposure to an acidic environment, such as lysosomes

(pH 4.5) (Albrecht et al. 2004). In general, most cells can

tolerate a certain amount of ROS, whereas higher levels of

ROS persist over a longer time and may result in cell

damage and subsequent induction of toxic effects (Wang

et al. 2007). Since the toxicity of the NPs is affected by the

level of induced ROS, the surface must be stable against

degradation to limit the quantity of free metal ions.

Potential (Eh)-pH diagram or Pourbaix diagram is

essential to investigate the thermodynamic of material cor-

rosion, by monitoring the regions of potential and pH where

the metal is: unreacted (region of immunity), protected by a

surface film of an oxide or a hydroxide (region of passivity)

or dissolved (region of corrosion) (McCafferty 2010). Fig-

ure 6 shows the Pourbaix diagram for both iron and palla-

dium elements in water containing fluoride ions (Villicana

et al. 2007). According to the diagram, iron will corrode and

produce Fe(II) and/or Fe(III) at potential zero and at pH

below 6, whereas palladium remains unreacted under these

conditions. This difference in stability is due to the higher

reactivity of iron towards oxidation (E�

FeðIIÞ=Fe = -0.44 V;

E�

FeðIIIÞ=Fe = -0.04 V), compared with palladium

(E�

PdðIIÞ=Pd = ?0.915 V). Moreover, iron forms a porous

oxide layer when exposed to water or air (Hill and Holman

2000), and consequently anodic (iron)/cathodic (iron oxide)

sites created at the surface trigger the process of corrosion.

The reactivity of iron towards oxidation reveals the

toxicity of uncoated IONPs (Pisanic et al. 2007), and

suggests more study into the biocompatibility of the coat-

ings on the long term (Hilger et al. 2003). The most

important source of toxicity of IONPs is described by

‘Fenton’ and ‘Fenton like’ reactions (Eqs. (16) and (17a,

17b), respectively) (Salgado et al. 2013), in which the

Fe(II) or Fe(III) reacts with H2O2 to produce ROS species.

Fe IIð Þ þ H2O2 ! Fe(III)þ HO� þ HO�

k ¼ 76M�1s�1ð16Þ

Fe IIIð Þ þ H2O2 ! Fe(II)þ HO�2 þ Hþ ð17aÞ

Fe IIIð Þ þ HO�2 ! Fe(II)þ O2 þ Hþ

k ¼ 0:01 M�1s�1ð17bÞ

Free ROS species exhibit a lack of specificity with

which they react (Pryor 1976), and this makes the study of

the oxidative mechanism in the toxicity of iron ions very

complex (Stohs and Bagchi 1995). However, ROS inter-

actions with biological components have been classified

into three types of reactions: electron transfer, radical

addition and atom abstraction, and identified to cause cell

damage (Moslen and Smith 1992).

The toxicity of Pd, Co pure metal and Pd43Co57 alloy

was tested in vitro for dental casting (Kawata et al. 1981).

It was shown that the cells multiply in the presence of Pd as

much as those for the control, and keep their natural form.

Whereas in the presence of Co, the cells degenerate with

time and approaches zero at 72 h of incubation due to the

cytoplasmic shrinkage and blister formation. On the other

hand, the binary alloy Pd43Co57 enhances the cells growth

and morphology compared with pure Co, showing a

monotonically increase of cell multiplication like the

control. These results indicate that the toxicity of Co may

be avoided when alloyed with Pd.

The corrosion of the binary alloy Pd80;8Co19;2 in syn-

thetic saliva (Goehlich and Marek 1990) produces a

selective dissolution of the less noble components ‘Co’ on

the surface of the alloy, leaving a Pd-enriched layer on the

surface. The results of corrosion are in accordance with

that of toxicity, the safety of a biomaterial largely depen-

dent on its corrosion resistance. Therefore, pure palladium

is non-toxic due to the low dissolution rate of palladium

ions (Wataha and Hanks 1996), while pure Co is not

stable and thus releases toxic cobalt ions.

Fig. 6 Pourbaix diagram showing iron and palladium species and

water stability region (reproduced from Villicana et al. 2007)

Prog Biomater (2016) 5:147–160 155

123

Despite belonging to essential trace elements of the

human body, the accumulation of cobalt ions is genotoxic

and may cause induce necrosis with inflammatory response

(Donaldsaon and Beyersmann 2005). The oral median

lethal dose (LD50) for soluble Co salts has been estimated

to be between 150 and 500 mg/kg body weight (Donald-

saon and Beyersmann 2005). Further, very low doses of Pd

are sufficient to cause allergic reactions in susceptible

individuals (Kielhorn et al. 2002). Oral LD50 of palladium

oxide is about 4.9 g/kg body weight (Nordberg et al. 2014).

Also high concentrations of Pd ions are capable of eliciting

a series of cytotoxic effects (Kielhorn et al. 2002).

Electrochemical corrosion test and immersion test were

performed at 37 �C for Pd93:85Co6:15 alloy sample (with a

density of 11.4 g/cm3) in mammalian Ringer’s solution

(Paulus et al. 1997). The tests results showed a long-term

corrosion rate of 7.7 9 10�8 lm/year, and a release of

0.7 lg/l of Pd(II) with 1.8 lg/l of Co(II) per year, indi-

cating a significantly high corrosion resistance of PdCo

compared with standard surgical implants (0.04 lm/year)

(Paulus et al. 1997).

According to the phase diagram of PdCo alloy (Fig. 7),

a single phase solid solution of substitutional Co atoms in a

Pd lattice is formed when the atomic percentage of Pd is

higher than 53 %. Consequently, the corrosion behavior of

the PdCo alloy will be similar to that of pure Pd. In fact,

palladium remains unreacted at normal pH or even acidic

environment, as stated in Pourbaix diagram (Fig. 6). Pure

palladium corrodes only in extremely acidic medium,

which is unlikely to occur in biological media. The selec-

tive dissolution of Co near or at the surface on the long-

term is possible (Paulus et al. 1997), and as a result Co-

depleted layer is formed. The alloy is then likely to exhibit

passivation behavior of pure palladium. An additional

dissolution of cobalt may occur by volume diffusion of

these less noble atoms to the surface (Pickering and

Wagner 1967).

Alloying Pd and Co not only induces ferromagnetism in

Pd atoms but also enhances the corrosion resistance of Co

atoms, which makes this alloy a good candidate for

biomedical application.

Very recently using magnetic nanoparticles for

enhancing the effectiveness of ultrasonic was shown

(Jozefczak et al. 2016). It was indicated that the effec-

tiveness of ultrasound (US) for medical applications can be

significantly improved by using sonosensitizers like mag-

netic nanoparticles with mean sizes of 10–300 nm. These

NPs can be more effectively heated because of additional

attenuation and scattering of US. Accordingly, this can

enhance the thermal effect of ultrasound (US) on the tissue

by increasing US absorption. The other interesting aspect

using magnetic NPs is that they are able to produce heat in

the alternating magnetic field (magnetic hyperthermia).

This is particularly important because it introduces syner-

getic application of ultrasonic and magnetic hyperthermia

which can lead to a promising treatment modality. In

particular, it was found that in the samples with magnetic

nanoparticles, the synergetic action of ultrasounds and

magnetic field allowed achieving better heating effect in

comparison to the heating by either US or alternating

magnetic field (AMF) alone. This synergistic effect was

confirmed by specific absorption rate (SAR) values. The

following SAR values were, respectively, obtained: 66

mW/g for magnetic hyperthermia, 175 mW/g for ultrasonic

hyperthermia, and 375 mW/g for both methods applied

simultaneously. This opens the ways to future potential

investigations of better utilization of NPs and ultrasound

for stand-alone magnetic hyperthermia therapy

applications.

Conclusions

Magnetic NPs are frequently employed in biomedical

research as drug delivery systems and/or magnetic reso-

nance contrast agents. Nevertheless, the safety issues of

these particles have not been completely solved because it

is difficult to compare the cytotoxicity data since the toxic

effects of NPs are influenced by many parameters (such as

size distribution, surface coating, magnetic properties, etc.)

(Auffan et al. 2006). Also, numerous studies showed con-

tradicting findings since different cell types will interact

with the same particle in different ways (Barua and Rege

2009). Moreover, the lack of coherence between various

research activities for establishing priorities among the

research needs is one reason why a toxicological profile of

these particles has not yet been well documented in the

literature. Therefore, along with the expanding applicationsFig. 7 Phase diagram of PdCo system obtained from FactSage

software (Bale et al. 2002)

156 Prog Biomater (2016) 5:147–160

123

of NPs and the growing numbers of consumer products

containing NPs, the release of these substances into the

environment is expected, and the impact of these materials

is increasing significantly (Zhu et al. 2012).

In this study, we have reviewed the basics of magnetic

properties and nanotoxicity of NPs for magnetic hyper-

thermia. Also, recent advances on the most used MNPs for

biomedical application were discussed. From this study, it

can be seen that despite its corrosion problem, iron oxide

NPs have received considerable attention. However, new

candidates such as PdCo NPs may have a great potential

for magnetic hyperthermia due to their high corrosion

resistance and good ferromagnetic behavior.

Some challenges need to be addressed on the design of

novel NPs, which must meet the demands of a particular

application. The elaboration of methods must be also sig-

nificantly improved to assess the toxicity of NPs, such as

reference biomaterials for safety testing. Synergetic

approaches combining magnetic and ultrasounds properties

must be also more investigated to improve the applicability

of magnetic NPs for magnetic hyperthermia therapy.

Acknowledgments The authors thank the Fonds Quebecois de la

Recherche sur la Nature et les Technologies (FQRNT) and the Fon-

dation Universitaire Pierre Arbour for their funding of this research.

Open Access This article is distributed under the terms of the

Creative Commons Attribution 4.0 International License (http://crea

tivecommons.org/licenses/by/4.0/), which permits unrestricted use,

distribution, and reproduction in any medium, provided you give

appropriate credit to the original author(s) and the source, provide a

link to the Creative Commons license, and indicate if changes were

made.

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